Multi-leaf radiation attenuator for radiation therapy with verification system

Abstract: A radiation therapy apparatus includes a compensator with moving leaves to attenuate a radiation beam and a compensator verification system to determine if the attenuating leaves are moving in accordance with desired position signals. Two radiation intensity monitoring chambers, one chamber on either side of a patient within a radiation beam, produce beam fluence data used by a computer to construct radiation absorption images and to evaluate the functionality of the compensator. The absorption image can be used for radiation dose verification and planning of suitable subsequent therapy sessions. The compensator may include a first plurality of radiation attenuating leaves and a second plurality of radiation attenuating leaves. The first plurality of leaves may be spaced with gaps between adjacent leaves within a radiation beam so that the leaves and the gaps there between divide the radiation beam into rays. The second plurality of leaves may be disposed directly under gaps in the first plurality of leaves so that each ray of beam can be occluded by a leaf in either the first or second plurality. The gaps eliminate interference between leaves; the staggering prevents radiation leakage between leaves. Each leaf, in both groups, may be moved between a first open state on one side of the beam, a closed state within the beam thus occluding one ray of the beam, and a second open state on the other side of the beam.

Description: This invention relates generally to radiation therapy equipment for the treatment of tumors, or the like, and specifically to a mechanism for regulating the dose of radiation within irregularly shaped zones within a patient and to a mechanism for verifying radiation intensity directed at, and the dose of radiation absorbed within, irregularly shaped zones of a patient. Medical equipment for radiation therapy treats tumorous tissue with high energy radiation. The dose and the placement of the dose must be accurately controlled to insure both that the tumor receives sufficient radiation to be destroyed, and that damage to surrounding and adjacent non-tumorous tissue is minimized. Internal-source radiation therapy places capsules of radioactive material inside the patient in proximity to the tumorous tissue. Dose and placement are accurately controlled by the physical positioning of the isotope. However, internal-source radiation therapy has the disadvantages of any surgically invasive procedure, including discomfort to the patient and risk of infection. External-source radiation therapy uses a radiation source that is external to the patient, typically either a radioisotope, such as 60 Co, or a high energy x-ray source, such as a linear accelerator. The external source produces a collimated beam directed into the patient to the tumor site. External-source radiation therapy avoids some of the problems of internal-source radiation therapy, but it undesirably and necessarily irradiates a significant volume of non-tumorous or healthy tissue in the path of the radiation beam along with the tumorous tissue. The adverse effect of irradiation of healthy tissue may be reduced, while maintaining a given dose of radiation in the tumorous tissue, by projecting the external radiation beam into the patient at a variety of "gantry" angles with the beams converging on the tumor site. The particular volume elements of healthy tissue, along the path of the radiation beam, change, reducing the total dose to each such element of healthy tissue during the entire treatment. The irradiation of healthy tissue also may be reduced by tightly collimating the radiation beam to the general cross section of the tumor taken perpendicular to the axis of the radiation beam. Numerous systems exist for producing such a circumferential collimation, some of which use multiple sliding shutters which, piecewise, may generate a radio-opaque mask of arbitrary outline. As part of collimating the beam to the outline of the tumor, the offset of the radiation beam, with respect to a radius line between the radiation source and the center of rotation of the radiation source, may be adjusted to allow the treated area to be other than at the center of rotation. Simultaneously changing the offset and the width of the radiation beam as a function of gantry angle allows tumorous tissue having an irregular cross-section within a plane parallel to the radiation beam to be accurately targeted. The width and offset of the radiation beam may be controlled by the use of a multiple-leaf circumferential collimator. Adjustment of the offset and size of the radiation beam at various gantry angles allows considerable latitude in controlling the dose. Nevertheless, even using these techniques, there is still a considerable amount of undesired dose imparted to healthy tissue, especially where a treatment volume is concave or highly irregular within the plane parallel to the radiation beam. A radiotherapy machine providing much reduced irradiation of healthy tissue is presented in u.s. Patent number 5,317,616 issued to Stuart Swerdloff et al, on May 31, 1994. The architecture described in the above application employs a radiation source constrained to rotate within a single gantry plane about a patient while intensity of individual rays of the radiation beam are modulated by a set of opaque leaves which move into and out of the radiation beam. The leaves slide into the radiation beam in a closed state and out of the radiation beam in an open state to allow unobstructed passage of a given ray of the beam. By employing appropriate planning techniques the dose absorbed by each slice of the tumor may be controlled to irradiate even tumors having a concave cross section within the gantry plane. This ability to control not just the outline of the radiation but the intensity of each individual ray allows extremely precise control of the irradiation volume. An entire tumorous volume may be treated by moving the patient with respect to the gantry plane and irradiating each tumor slice separately. The compensator and gantry configuration together substantially increase the ability to conform a radiation dose to an arbitrarily shaped tumor while employing a simplified two-dimensional protocol. Besides simplifying the irradiation protocol, a single plane configuration provides a number of other benefits. These benefits include the ability to use a single ring gantry to support both a radiation source and a CT system, the advantage of reduced interference between the radiation source and patient (or table) and the advantage of simplified shielding requirements. Furthermore, the simplified architecture enables a therapist to employ a helical scanning method to smoothly irradiate along the length of a tumor so as to avoid irradiation hot spots or irradiation gaps. Despite the advantages of constraining a therapy machine to operation with a single plane, a single plane machine presents various problems. First, there is the need for dose verification. The destructive potential of a radiation beam to healthy tissue and the necessity of insuring that tumorous tissue receives sufficient radiation makes treatment verification a required part of radiation therapy. With conventional therapy machines, films may be exposed during a therapy session both to confirm the location of an irradiated area and to provide a record of radiation dose. Because the radiation source employed by the above described gantry configuration constantly sweeps around a gantry to produce a moving beam, a film verification system employing traditional film movement techniques would be unworkable. The verification problem is more profound if scanning is performed in a helical manner. In addition, a single plane system must employ a compensator capable of varying the intensity of individual rays of a beam in order to properly treat a tumor. The reliability of such a compensator must be extremely high even a single improperly attenuated beam ray passing through the compensator undetected and irradiating healthy tissue due to a failed compensator component could result in severe damage to healthy tissue. In practice, regardless of accurate machining techniques, small gaps between the leaves in the compensator described above, that allow unattenuated rays to pass, are needed to prevent frictional contact. The exacting tolerances necessary to minimize the size of the leaf gaps are costly and such high tolerance components are prone to failure. Another problem with the compensator described above is that the rays are not uniformly attenuated by the moving leaves because the leaves cannot be moved instantaneously. A leaf initially occludes the entire depth of its associated ray within the beam. As the leaf begins to move out of the beam, part of the ray is occluded and another part is left unobstructed. Eventually, the entire ray is unobstructed. The same non-uniform beam attenuation is again encountered when the leaf moves back into the beam. This gradation in attenuation can be minimized by equipping the compensator with more powerful actuators to drive the leaves in and out of the fan beam width more rapidly. Bigger actuators, however, are more costly to utilize and maintain. Alternatively thinner leaves may be used which are light weight so that they can be moved more quickly. This, however, demands more actuators, creates more leaf gaps, and must be accommodated by a much more complicated control system. US4,987,309 discloses a radiation therapy machine in which a multileaf collimator defines a cross section of a fan beam relatively precisely but does not discuss controlling ray intensity. Control of the multileaf collimator is not discussed in detail. The present invention provides a radiation therapy machine having a radiation source for producing a radiation beam directed toward a patient at a gantry angle, the beam including a plurality of adjacent rays, the machine having an attenuating means disposed between the radiation source and the patient, characterised in that the attenuating means is capable of independently controlling the fluence of each ray of the beam, and in that the machine further comprises a verification system comprising: a pre-patient monitor generally disposed between the attenuating means and the patient for determining a pre-patient fluence of each ray of the beam prior to entering the patient; and a post-patient monitor generally disposed opposite the pre-patient monitor with respect to the patent and within the fan beam for determining a post-patient fluence of each ray of the beam exiting the patient. In one embodiment, the attenuating means is a plurality of radiation attenuating leaves in a support structure positioned generally between the radiation source and the patient. The support structure guides the leaves between a closed state within the radiation beam, each leaf thus occluding one ray of the beam, and an open state outside of the radiation beam to allow unobstructed passage of the ray. A motivator independently moves each leaf between the open and closed states to effect an open-to-closed desired ratio producing the fluence of each ray. The motivator may be a first set of actuators connected by linkages to individual leaves to move the leaves with the movement of the armatures. A position sensor determines when each leaf is in the open state and when each leaf is in the closed state and produced an actual ratio of the period of time the leaf is in the open state to the period of time the leaf is in the closed state. An error detector generates an error signal by comparing the actual ratio to the desired ratio. The present invention may also provide a mechanical compensator system with extremely high reliability. An error signal generated by an error detector may be used to indicate a malfunction. In another embodiment, the supporting structure may guide the leaves between a closed state centered within the radiation beam, the leaf thus occluding one ray of the beam, a first open state with the leaf displaced outside of the radiation beam on a first side of the one ray, and a second open state with the leaf displaced outside of the radiation beam on a second side of the one ray. In this embodiment the motivation means moves each leaf alternatively between the first open and closed state and between the closed and second open states. The present invention may include a compensator that attenuates individual rays of a high energy fan beam to provide uniform attenuation across the thickness of the fan beam. Provided the acceleration and velocity of the lead in each direction is identical, the uneven radiation caused by motion of the leaf in one direction is cancelled by the radiation caused by the motion of the lead in the second direction. In yet another embodiment, the plurality of attenuating leaves in the attenuating means above includes a first plurality of radiation attenuating leaves second plurality of radiation attenuating leaves and the supporting structure includes a first supporting structure for guiding the first plurality of leaves between a closed state, each leaf thus occluding every other ray of the beam, and an open state outside of the radiation beam and a second support structure for guiding the second plurality of leaves between a closed state, each leaf thus occluding those rays of the beam not occluded by the first plurality of leaves when the latter are in the closed state, and an open state of the radiation beam wherein the first plurality of leaves in the closed state are positioned closer to the radiation source than the second plurality of leaves in the closed state. It is thus yet another object of the invention to provide a compensator that eliminates the possibility of radiation leakage between the leaves of the compensator without requiring high tolerance machining or risking the possibility of interference between adjacent leaves. The use of two levels increases the spacing between the leaves eliminating their interference yet allows a slight overlap between the leaves of the two levels to eliminate passage of unintended radiation. A post-patient monitor may be disposed opposite the pre-patient monitor with respect to the patient and within the fan beam for determining a post-patient fluence of each ray of the beam exiting the patient. An absorption calculator compares the pre-patient fluence to the post-patient fluence to produce an absorption value for each ray, the absorption values together providing an absorption profile for the fan beam at a given gantry angle. Yet another object of the invention is to provide a record of radiation dose distributed to a tumor site. This dose record may be used by a radiotherapist to verify radiation dose throughout the volume of a tumor. In addition, the dose record reduces the possibility that interruptions in a therapy session will require termination of the session. Sessions can resume where they were interrupted. Furthermore, the dose record may be used in planning suitable subsequent therapy sessions. The foregoing and other objects and advantages of the invention will appear from the following description. In the description, reference is made to the accompanying drawings which form a part hereof and in which there is shown by way of illustration several preferred embodiments of the invention. Such embodiments do not necessarily represent the full scope of the invention, however, and reference must be made therefore to the claims herein for interpreting the scope of the invention. In the drawings Fig. 1 is a perspective view of the compensator assembly used in the present invention, showing the compensator leaves and their associated electromagnetic actuators; Fig. 2 is a cross section of the compensator assembly of Fig. 1 along line 2-2 showing the trapezoidal aspect of each compensator leaf, for a fan beam of radiation, and the guide rails for supporting the compensator leaves when they move; Fig. 3 is a cutaway perspective view of a set of guide rails and one leaf of Fig. 2 showing a collar for supporting the leaf; Fig. 4 is a plan view of a portion of the mounting rack and one leaf of Fig. 1 showing a trigger bore, a trigger and a light emitting diode and light detector pair with the leaf in its fully closed position; Fig. 5 is a block diagram showing the elements of a radiation therapy apparatus incorporating a conventional CT scanner and the compensator of the present invention and including a computer suitable for controlling that compensator per the present invention; Figs. 6(a)-(d) are dose distributions of a hypothetical tumorous region showing dose intensity by lines of equal dose, with Fig. 6(a) showing a desired dose distribution and Figs. 6(b), (c), and (d) showing progressive actual dose distributions after two, three and ten iterations per present invention; Fig. 7 is a diagrammatic representation of a patient receiving radiation therapy, showing the scatter kernel and the coordinate system used to describe the present invention; Fig. 8 is a perspective representation of a monodirectional scatter kernel associated with a radiation beam at one gantry angle; Fig. 9 is a perspective representation of a composite multidirectional scatter kernel associated with a plurality of radiation beams at multiple gantry angles; Fig. 10 is a block diagram depicting the fluence profile calculator which takes a desired dose map and calculates a fluence profile; Fig. 11 is a block diagram depicting the overall iterative method of controlling the compensator of the present invention, employing the fluence profile calculation method of Fig. 10; Figs. 12(a)-(c) are perspective views of plots showing the error between the desired dose distribution and the actual dose distribution obtained with the present invention for one, two and four steps of iteration respectively. Figs. 13(a)-(c) are schematic views showing the relationship between an irradiation window and adjacent tumor slices as the radiation source rotates about the gantry from 0° to 90° to 180°. Fig. 14 is a schematic view showing the general orientation of the monitoring chambers in relation to the compensator and the patient. Fig. 15 is a block diagram depicting the fluence modulation method of changing leaf ratios to adjust ray fluences. Fig. 16 is a perspective view of a second embodiment of assembly used in the present invention showing two levels of the compensator leaves, their associate pneumatic cylinders and their associated solenoid assemblies; Fig. 17 is a cross-section of the compensator assembly of Fig. 16 along line 17-17 showing both levels of compensator leaves, the trapezoidal aspect of each compensator leaf, the associated solenoid stopping assemblies, and the guide rails for supporting the compensator leaves when they move; Fig. 18 is a cutaway perspective view of a set of guide rails and one leaf from the first level and one leaf from the second level of Fig. 16; Figs. 19(a)-(c) are side views of a leaf showing the relationship between a leaf, the beam and the solenoid assemblies when the leaf is in the first open Fig. 19(a), closed Fig. 19(b) and second open Fig. 19(c) states; Fig. 20 is a detailed cross sectional view of the solenoid stop assembly shown in Fig. 17 taken along the line 20-20; Fig. 21(a)-(g) are graphs showing the changing fluence gradient across the beam as a leaf is moved from a closed state in Fig. 21(a) into a first open state in Fig. 21(c), back into the closed state in Fig. 21(e), into a second open state and again back into the closed state Fig. 21(g). Referring to Fig. 1, a radiation therapy unit 10 suitable for use with the present invention includes a radiation source 12 producing a generally conical radiation beam 14' emanating from a focal spot 18 and directed towards a patient 17 (not shown in Fig. 1). The conical beam 14' is collimated by a radiation opaque mask 16 constructed of a set of rectangular collimator blades to form a generally planar fan beam 14 centered about a fan beam plane 20. I. The Compensator A compensator 22 is centered in the fan beam 14 and about the fan beam plane 20, prior to the radiation being received by the patient 17, and includes a plurality of adjacent trapezoidal leaves 30 which together form an arc of constant radius about the focal spot 18. The leaves 30 are held in sleeves 24. The sleeves 24 are constructed of radio-translucent materials and attached at their inner ends 23 to a mounting plate 26 which is fixed relative to the focal spot 18. The mounting plate 26 is constructed of a sturdy, radiopaque material and is positioned just outside the fan beam 14 to prevent interference with the fan beam 14. Preferably, the leaves 30 of the compensator 22 subtend the entire fan beam 14 to divide the fan beam 14 into a set of adjacent slab-like rays 28 at offset angles 1. Referring also to Fig. 2, each sleeve 24 is open at its outer end 27 to receive, by sliding, a comparably sized trapezoidal leaf 30 constructed of a dense, radiopaque material such as lead, tungsten, cerium, tantalum or a related alloy. Each leaf 30 may slide completely within its corresponding sleeve 24 to block the ray 28 associated with that sleeve 24. When the leaf 30 blocks its corresponding ray 28, it is referred to as being in a "closed state". The sleeves 24 are of ample length to permit each leaf 30 to slide out of the path of the fan beam 14, so as to leave its corresponding ray 28 completely unobstructed, and yet to still be guided by the sleeve 24. In this non-blocking position, a leaf is referred to as being in the "open state". Each leaf 30 moves rapidly between its open and closed states by means of a primary corresponding relay-like electromagnetic actuator 32 connected to the leaf 30 by a slider member 34. The actuators 32 have internal armatures (not shown) received within solenoid electromagnets. The armature may be moved at high velocity by means of varying electrical excitations of their associated electromagnets. The electrical excitations are provided by a compensator control (not shown in Figs. 1 or 2) to be described below. The actuators 32 are capable of applying high forces to the leaves 30 to move them rapidly and independently between the open and closed states. Each leaf 30 is also provided with a back up actuator 35 located below its primary actuator 32 on the outer edge of the leaf 30. The secondary actuator 35 is employed when the primary actuator 32 fails as will be described in more detail below. Referring now to Figs. 2 and 3, the leaves 30 are supported and guided within the sleeves 24 by guide rails 36 fitted into notches 38 cut along the edges of the leaves 30. The notches 38 allow the guide rails 36 to slidably retain the leaves 30 within the sleeves 24 during motion between the open and closed states. In the closed state, the inner end 40 of each leaf 30 is captured by a rigid collar 42 attached to the mounting plate, which aligns the leaf 30, more accurately than may be done by the guide rails 36, with the mounting plate 26 and hence with the fan beam 14. Whereas the guide rails 36, which are ideally radio-translucent, are relatively insubstantial, in contrast, the collar 42, positioned outside the fan beam 14 on the mounting plate 26, need not be radio-translucent and hence is more substantial in construction. A collar (not shown) similar to collar 42, supports each leaf 30 when it is fully in the open state. Because the leaves 30 spend most of their time fully in the open or closed states, they are, at most times, firmly located by a supporting collar 42. Of concern is the reliability of leaf 30 switching because one inoperable leaf 30 would prevent treatment. The number of switching cycles S (opening and closing) in a year is given by the following equation: S =250 ρ LB Δ z where 250 is typical of the number of treatment days per year in North America and P is the number of patients treated per day. S is going to be on the order of millions of cycles per year. If the probability of a single leaf 30 failing in a year is P 1 then the probability of any leaf within the system failing P is given by: P =1- (1- P 1 ) " where n is the number of leaves 30. Since the leaves 30 are light, the slide member 34 is under little stress and so it has a low probability P slide of failure. Referring now to Figs. 3 and 4, a light opaque trigger member 19 is integrally attached to the inner edge 40 of each leaf 30, the trigger member 19 extending laterally outward parallel to the movement axis 25 of each leaf 30. When the leaf 30 is in the closed state (see Fig. 4), the trigger member 19 passes through a trigger bore 31 in the back wall 29 of the mounting plate 26. A plurality of light emitting diode 21 and light detector 15 pairs are positioned on the outer surface of the back wall 29, the elements of each pair opposing each other on opposite sides of an associated trigger bore 31. When a leaf 30 is in the closed state, its associated trigger member 19 extends through the trigger bore 31 and blocks the light path 33 between the light emitting diode 21 and light detector 15. Referring to Fig. 3, when the leaf 30 is not in the closed state, the trigger member 19 does not extend through the trigger bore 31 and hence light passes from the emitting diode 21 to the light detector 15. The probability of the light detector 15 failing is P verify . An axially compressible spring 39 is provided between the outer edge 41 of each leaf 30 and the front wall (not shown) of the mounting plate 26 to move the leaf 30 to the closed position absent force from its actuator 32. A system employing a light detector 15, primary and backup actuators 32, 35 and which are regularly serviced, would have downtime due to failure of a leaf 30 according to the probability equation: P 1 =1- (1- P slide ) (1- P verity ) [1-( P actuator N maint ) 2 ] wherein P actuator is the probability of the actuator or control electronics failing during a year and N maint is the number of maintenance services in the year (e.g. it could be every day). The exponent 2 accounts for the failure of both actuators in the maintenance period. This equation basically says that failure of a leaf 30 will happen if the slider 34 fails, the verification system fails or both of the actuators 32, 35 fail before the system can be serviced. Referring to Figs. 5 and 14, a pre-patient multi-segment ion chamber 47 is positioned between the compensator 22 and the patient 17. Each ray 28 of the fan beam 14 subtends a separate monitor segment 49 as it passes through the ion chamber 47. A post-patient multi-segment ion chamber 53 is positioned directly opposite the radiation source 12 on the gantry 44 so as to intercept the fan beam 14 as it exits the patient 17. The separate monitor segments 54 of the second ion chamber 53, like the monitor segments 49 of the first ion chamber 47, are each subtended by individual rays 28 of the fan beam 14. The ion chambers 47, 53 produce signals indicating the fluence of rays 28 (as generally understood in the art) and as used by the computer 51 to determine radiation dose in a manner to be described below. II. Radiation Therapy Hardware Referring now to Pig. 5, the radiation source 12 is mounted on a gantry 44, the latter rotating within the fan beam plane 20 about a center of rotation 45 in the patient 17 so that the fan beam 14 may irradiate a slice of the patient 17 from a variety of gantry angles . The radiation source 12 is controlled by a radiation control module 48 which turns the radiation beam 14 on or off under the control of a computer 51. A compensator control 52 directed by a timer generating desired position signals provides electrical excitation to each electromagnet to control, separately, the actuators 32 to move each of the leaves 30 in and out of its corresponding sleeve 24 and ray 28 (see also Fig. 1). The compensator control 52 moves the leaves 30 of the compensator 22 rapidly between their open and closed states to either fully attenuate or provides no attenuation to each ray 28. Gradations in the fluence of each ray, as needed for each fluence profile, are obtained by adjusting the relative duration during which each leaf 30 is in the closed position compared to the relative duration during which each leaf 30 is in the open position, for each gantry angle. The ratio between the closed and open states or the "duty cycle" for each leaf 30 affects the total energy passed by a given leaf 30 at each gantry angle and thus controls the average fluence of each ray 28. The ability to control the average fluence at each gantry angle permits accurate control of the dose provided by the radiation beam 14 through the irradiated volume of the patient 17 by therapy planning methods to be described below. The compensator control 52 also connects with computer 51 to allow program control of the compensator 22 to be described. A tomographic imaging system 11 employing an x-ray source 46 and an opposed detector array 50 may be advantageously mounted on the same gantry 44 as the radiation source 12 to produce a tomographic or slice image of the irradiated slice of the patient 17 prior to radiation therapy for planning purposes. Alternatively, such tomographic imaging may be performed on a separate machine and the slices aligned according to fiducial points on the patient 17. A gantry control module 9 provides the signals necessary to rotate the gantry 44 and hence to change the position of the radiation source 12 and the angle  of the fan beam 14 for the radiation therapy, as well as for the computed tomography x-ray source 46 and detector array 50 also attached to gantry 44. Gantry control module 9 connects with computer 51 so that the gantry may be rotated under computer control and also to provide the computer 51 with a signal indicating the gantry angle  to assist in that control. Control modules for the tomographic imaging system 11 include: x-ray control module 56 for turning on and off the x-ray source 46, and data acquisition system 58 for receiving data from the detector array 50 in order to construct a tomographic image. An image reconstructor 60 receives the data from the data acquisition system 58 in order to assist in "reconstructing" a tomographic treatment image from such data according to methods well known in the art. The image reconstructor 60 also communicates with computer 51 to assist in high speed computations used in the present invention as will be described below. The tomographic treatment image allows verification of the patient setup just prior to radiation therapy treatment. An image reconstructor 60 typically comprising a high speed array processor or the like may use the actual fluence signals 57 and barrier signals 59 to produce a tomographic absorption image to be used for verification and future therapy planning purposes as described in more detail below. A terminal 62 comprising a keyboard and display unit 63 allows an operator to input programs and data to the computer 51 and to control the radiation therapy and tomographic imaging equipment 10 and 11 and to display tomographic images produced by the image reconstructor 60 on the display 63. A mass storage system 64, being either a magnetic disk unit or tape drive, allows the storage of data collected by the tomographic imaging system 11 and the multi-segment ion chambers 47, 53 for later use. Computer programs for operating the radiation therapy system 10 will generally be stored in mass storage unit 64 and loaded into the internal memory of the computer 51 for rapid processing during use of the system 10. During operation of the radiation therapy unit 10, the compensator control 52 receives from the computer 51 a fluence profile for each gantry angle. The fluence profile describes the intensity or fluence of each ray 28 of the radiation beam 14 that is desired for that gantry angle  at a given position of the patient support table (not shown) as translated through the radiation beam 14. The collection of fluence profiles over a range of rotation gantry angles is termed a "treatment sinogram". III. Therapy Planning Software The generation of a treatment sinogram needed to obtain the full benefits of the above described compensator is performed by specially developed software running on the computer 51 and reconstructor 60. Although the treatment planning is performed in software, it will be recognized that the planning may also be implemented in discrete electronic circuitry dedicated to this operation and that such dedicated circuitry may be employed to provide even greater speed to this process. Referring to Fig. 6(a), the generation of the desired treatment sinogram to control compensator 22 begins with the definition of a desired dose map 66. A typical desired dose map 66 assigns a relatively high radiation dose, within a dose constraint, to an area of tumorous tissue 68 and a second lower radiation dose to the area of healthy tissue 70 outside of that region. The healthy tissue 70 may include an area 72 including a radiation sensitive organ or the like to which an even lower radiation dose may be assigned. The desired dose map 66 is stored within the memory of computer 51 as an array of elements each element holding one digital value, and may be most easily entered by displaying the tomographic view of the slice of patient 17 on the display 63 of the terminal 62 and manually tracing around the tumorous area 68 using a track-ball or a similar input device as is well understood in the art. Standard area-filling computer programs may be used to transfer the dose values assigned to each traced region to the appropriate element in the array of memory representing the desired dose map 65. Each element of the dose map 66 thus defines the dose desired at each of the plurality of volume elements 74 ("voxels") within a slice of the patient 17. Referring to Fig. 7, each voxel 74 of the patient 17 may be identified by a vector r defined from a given reference point 76. The dose at each voxel 74 is D( r ). A. Converting Dose to Terma 1. Terma Generally, the dose at any voxel r will depend on the energy received at that voxel r from radiation scattered from adjacent voxels (where adjacent voxels r include the voxel r , i.e., the radiation received directly from the radiation source 12). The dose D( r ) for a given voxel r is given by the following formula: D ( r ) = ∫ T ( r ') A ( r - r ') d 3 r ' where T( r ') is a value indicating the magnitude of the primary total energy released from r ' per unit mass of that voxel r ' and is called the "terma" (total energy released per unit mass). For a monoenergetic external radiation source, the terma rate T ˙ ( r ') is described by: where µ / ρ is an effective mass attenuation value at the voxel r , E is the energy of the radiation photons in Joules,  ˙ is the distribution of the fluence rate (flux density). The integration of energy times fluence rate over time is energy fluence Ψ( r ) where: hence T ( r ') = µ ρ ( r ')Ψ( r ') Equation (4) basically relates how much energy from the ray 47 interacts with the voxel r'. 2. Convolution Kernel A( r - r '4) is a convolution kernel describing non-stochastic energy transport or scattering in a uniform medium. A( r - r ') thus describes how the energy from each voxel r ' spreads to contribute to the dose at voxel r . The kernel A( r - r ') may be generated using a Monte Carlo method as is generally understood in the art. As mentioned, it is a three-dimensional function indicating the fraction of energy absorbed at voxel r per unit of energy released at voxel r '. The energy emitted from the terma of each voxel r ' finds its source in a directed ray 47 from external radiation source 12 and thus A( r - r ') is generally anisotropic as suggested in Fig. 8, spreading outward away from the entry of ray 28. Energy conservation requires that: ∫ A ( r '-) d 3 r ' = 1.0 That is, if the energy transferred by the primary interaction were all deposited on the interaction point, the kernel would be approximated as a delta function. Referring still to Fig. 8, the anisotropy of A( r - r ') is related to the gantry angle  and thus of the angle of incidence of the ray 28 at r '. If the gantry angles  at which the patient 17 is irradiated are predetermined, a multi-direction convolution kernel B( r - r '), shown in Fig. 9, may be created-from weighted superimposition of the kernels A( r - r '). Referring to Fig. 9, assuming that the spreading of radiation is approximately equal for all beam directions and the rays 28 from each gantry angle  contribute equally to the terma at voxel r ', then the multidirectional convolution kernel reduces to a "isotropic" form as follows: where n is the number of discrete gantry angles from which rays 28 are projected. For multiple rays 28 at different gantry angles, the total dose at a given voxel r is the sum of doses from each constituent beam, therefore: D ( r ) = ∫ T ( r ') B ( r - r ') d 3 r ' where T ( r ') = nT ( r ') i , the latter term being the contributed portion of the terma for the ith gantry angle. This simplification assumes that the contribution to the terma from each ray 28 is equivalent and takes advantage of the distributive property of convolution. Errors in this assumption are reduced by the filtration to be discussed later. Equation (7) substantially simplifies the calculation of dose from terma but still requires a convolution for each voxel r times the total number of voxels r ' to calculate the dose over the entire patient volume. Preferably, therefore, the calculational efficiency of the fast Fourier transform can be used and equation (7) converted to the following: D ( r ) = F -1 { F { T ( r ')· F { B ( r - r ')}} where F and F -1 symbolize Fourier and inverse Fourier transforms respectively. This simplification of equation (8) requires that the kernel B( r - r ') be spatially invariant and relies on the convolution theorem which states that convolution of two spatially invariant quantities in a space domain is equivalent to multiplication in the frequency domain. The assumption of the spatial invariance of B( r - r ') is correct only to a first order approximation. Typically, the kernel B( r - r ') for an external radiation source 12 is a complex function of: (1) beam hardening of a polyenergetic x-ray beam (i.e., the effect of the filtration of the patient 17 in increasing the proportion of high frequency or high energy radiation components), (2) the number of rays 28 crossing each voxel, and (3) exponential attenuation by the patient mass. In the preferred embodiment, this first factor, beam hardening, is neglected because it is an effect smaller than the attenuation problem. Thus, the photon energy spectrum in the patient 17 may be assumed to be the same as that of the external radiation source 12. This simplification is not required, however, and it will be understood that beam hardening could be accurately accounted for by representing a photon energy spectrum by a finite number of separately convolved energy intervals. The second factor, the difference in the number and orientation of rays 28 that cross each voxel, caused by the geometry of a finite number of gantry angles and the fan orientation of the beam 14, affect spatial invariance. Problems arising from the fan orientation of the beam (in contrast to a parallel beam geometry) are largely solved if there is a full rotation of the gantry 44. Errors resulting from the fact that irradiation is performed at only a finite number of gantry angles have been determined to be acceptable. The third factor affecting the assumption of spatial invariance is the attenuation of the medium. This affects the fractional contribution of the total terma from the beams at each gantry angle. Accordingly, in those steps of the planning procedure, as will be noted below, where accurate calculation of dose is critical, the dose distribution is calculated separately for each beam based on the attenuation of overlying voxels, such attenuation deduced from the parameters of the tomographic image. In this case the simplification of equation (8) may not be employed and repeated convolutions must be performed. In certain steps in the planning process, however, as will be noted, an estimate is sufficient and in these cases B( r - r ') is assumed to be spatially invariant and the dose calculated according to equation (8). Production of terma values from a desired dose map 75 is then simply the process of inverting equation (8) as follows: This inversion requires that there be no significant "zeros" (typically at high frequencies) in the denominator term F{B( r - r ')}or more simply that the kernel B( r - r ') be spatially compact (i.e., the Fourier transform of a spatially compact kernel will have significant high frequency content). It has been determined by the present inventors that the kernels dictated for patients 59 are sufficiently compact to allow this Fourier deconvolution. Referring now to Fig. 10, this deconvolution to produce a terma map 82, giving the terma for each voxel r , from the desired dose map 66, is represented by process block 80. B. Converting Terma to Voxel Energy Fluence Knowing the terma map 82, the energy fluence Ψ( r ') , which is a measure of the beam intensity, can be determined at each corresponding voxel by equation (4) from a knowledge of m/r as follows: Ψ( r ') = T ( r ') µ ρ ( r ') The value µ / ρ of may be estimated and considered a constant or actual m/r may be deduced from the tomographic scan data collected by means of the tomographic imaging system 60, (shown in Fig. 5). In this manner and as illustrated by process block 84 of Pig. 10, a fluence map 86 giving the fluence at each point of the terma map may be determined. C. Converting Voxel Energy Fluence to Energy Fluence Profile The energy fluence Ψ( r ') at each voxel r ' is related to the energy of the ray 28 exiting the compensator 22 by the relation: where Ψ o (,) is the energy fluence for a given ray 28 as described by δ(ρ- r · r ) at the exit of the compensator 22 and serves to define the fluence profile of the compensator and  and  are the gantry angle and the offset angles of the ray 28 as previously described. The exponential term represents the attenuation of the ray 28 from the exit of the compensator 22 to the voxel r caused by the mass of the patient 17 where by µ / ρ ( r ) is the attenuation for each voxel r along the ray 28, ρ( r ) is the density of each voxel r , SSD(,) is the distance between the exit of the compensator 22 and the surface of the patient 17, r is a unit vector along r (where the origin is now assumed to be the center of rotation of the gantry 44), and p is the perpendicular distance from the gantry's center of rotation 45 and the ray 28. The vector is simply a vector along the ray 28 to provide an integration variable. The fluence at each voxel r is related to the fluence of the radiation beam 14 emitted from the compensator 22 by equation (11). In the preferred embodiment, the density and attenuation of each voxel r , µ / ρ ( r ) and ρ( r ) are assumed to be constant and the fan beam of radiation is approximated by a parallel beam, hence SSD 2 (,) / | t | 2 = 1. Borrowing from the mathematics of tomographic image reconstruction, the fluence map 86 may be "reverse" back projected (i.e. projected) by projector 85 to determine a fluence profile to be produced by the external-source necessary to generate the desired fluence map and hence dose. , This projection is simply the opposite of a typical back projection used to form an image of a tomographic slice of a patient 17 from a series of projections taken in a tomographic imaging system. Because a projection is a line integral across a distribution, the energy fluence distribution for each voxel (equation (11)) is first differentiated with respect to the rayline t : The line integral of d ψ( r ) / dt along t , corrected for attenuation and inverse-square falloff, then represents the projection operation and ψ o (,) the fluence profile over the offset angles  of each gantry angle , is: - (ρ- r - r ) d t The projection of equation (13) is represented by projector 85 in Fig. 10. The projection process, for the purpose of computing fluence profiles for the compensator 22, differs in a fundamental way from the simple inverse of tomographic back projection. The difference is primarily in a concern for the sharpness in the transition of the dose between the irradiated tumorous tissue 68 and the healthy tissue 70. Sharpness in this transition region reduces the irradiation of healthy tissue 70 and is preferred over actual fidelity of the dose to the desired dose map 66. For this reason, the fluence map 86 from the fluence calculator 84 is prefiltered as shown by process block 88 to produce a filtered fluence map ψ'(,) as follows: ψ'(,) = F -1 { F {ψ(,)|ω|}. where ψ(,) is the fluence map 86 and |ω| is a ramp filter in frequency space and the '+' subscript indicates the positive component of the filtering result. This prefilter 88 serves to increase the high frequency content of the fluence map 86 and thus to aid in rapid transition of dose at the tumor/non-tumor interface. It is noted that this prefilter 88 is similar to the filter used in tomographic imaging's "filtered" back projection. That is, like tomographic imaging, the filter de-emphasizes the low frequency components of the projection in producing image data. In addition other prefilters may be applied to correct for the use of the radially symmetric kernel (equation (6)) in computing the dose map from the terma map composed from the fluence profile. In practice the computation of a terma map, the generation of a fluence map and the calculation of the fluence profiles need not be performed as discrete steps but may be accomplished by a direct projection of the dose map with appropriate filtering. The filtering is accomplished by a "fast inversion filter" which combines in projection space the operation of deconvolution and ramp filtration. This may be symbolically specified by the following equation ψ(,) = ρ{ D ( r )} ⊗ I ( t ) where ρ refers to a projection operation and I(t) is the fast inversion filter. The ⊗ operators refers to a convolution operation such as would normally be done in Fourier space using a fast Fourier transformation. Referring still to Fig. 10, the fluence profile calculations of block 78, including the deconvolver 80, the fluence calculator 84, the prefilter 88 which includes any projection space filter (such as a ramp filter, a fast inversion filter followed by truncation of zeros), and the projector 85 thus produce fluence profiles which together create an estimated treatment sinogram 87' from the desired dose map 66. The fluence profile calculator 78 may use the Fourier convolution of equation (9) in the estimate of the fluence profiles at this stage, accepting minor inaccuracies in that process, to be corrected at a later stage, as will be described below. D. Iteration Referring now to Fig. 11, the fluence profile calculator 78 converts the desired dose map 66 to an estimated treatment sinogram 87', however the fluence profiles of this estimated treatment sinogram 87' may not be used to control the compensator 22 because, in general, the estimated treatment sinogram 87' will include positive and negative values of fluence. Only positive values of fluence are physically realizable by the compensator 22; a negative value of fluence would represent a ray 28 that absorbed radiation along its path which is physically unrealizable. Accordingly, at process block 88, the fluence values of the estimated treatment sinogram 87' are truncated to positive fluence values 89. As a result of this truncation, the estimated treatment sinogram 87' no longer produces the desired dose map. The amount of error resulting from the truncation producing the positive fluence profiles 89 is determined by back-projecting the positive fluence values 89 to an actual dose map, 90 deviating from the desired dose map 66. This back projection is accomplished by computing a fluence map from the positive fluence values 89 per equation (11) and a terma map per equation (4) and then convolving the terma map with the kernel per equation (7) to establish the actual dose map 90 per process block 92 of Fig. 11. In this back projection, the assumption of spatial invariance of the convolution kernel B( r - r ') is not made so as to produce a more accurate actual dose map 90. The projection of a fluence profile onto a patient 17 to compute a dose map may be performed by a number of other procedures known to those of ordinary skill in the art. The actual dose map 90 is compared to the desired dose map 66 to produce residual dose map 96 as indicated by process block 94. In the preferred embodiment, the comparison subtracts from the values of each voxel r of the actual dose map 90, the greater of: a) the corresponding value of desired dose map 66, or b) a predetermined upper dose constraint. The predetermined upper dose constraint is a threshold number that is deemed an acceptable dose to tumorous tissue 76. Clearly, other methods of quantifying the difference between the desired dose map and the actual dose map will be apparent, from this description, to those of ordinary skill in the art. The result of this comparison process 94 is to produce a residual dose map 96 shown in Fig 12(a). This residual dose map 96 is then, again, operated on by the fluence profile calculator 78 (in lieu of the desired dose map 66) to produce an error fluence profile 98 (in lieu of the estimated treatment sinogram 87). A thus produced error fluence profile 98 is subtracted by subtracter 100 from the estimated treatment sinogram 87' to produce a new estimated treatment sinogram 90. As shown in Fig. 11, this new estimated treatment sinogram 87 is repeatedly operated on by process blocks 88, 92, 94 and 78 for a predetermined number of iterations, the magnitude of the error fluence profile 98 values decreasing with each iteration as shown in Figs. 12(b)-(c) until a suitably low error fluence profile 98 is obtained. The new estimated fluence profile 87 is then truncated per process block 88 to produce a final sinogram 91 for use in controlling the compensator 22, as previously described. Referring again to Figs. 6(b), (c) and (d), dose maps obtained by the present invention corresponding to a desired dose map 66 of Fig. 6(a) are shown after: one iteration (Fig. 6(b)); two iterations (Fig. 6(c)); and ten iterations (Fig. 6(d)). The variations in dose in the target volume shown in Fig. 6(d) are plus or minus 2% about the predetermined upper limit of 1,000 Cgy. E. Converting Slice Data To Helical Data In order to eliminate radiation hot spots and radiation gaps along the length of a tumor, helical scanning, in which the translation table is continuously moved along a "z axis" through the gantry 44 as the gantry 44 rotates so that the radiation fan beam 14 sweeps a helical pattern through the tumor site, is desirable. Helical scanners also reduce irradiation time because the start and stop motion of the translation table is eliminated. Nevertheless, because of the constant translation of the patient during helical scanning, the treatment sinogram, as discussed above, must be modified. Referring to Figs. 13(a)-(c), a simplified cylindrical tumor 97 with its axis coincident with the axis of rotation 45 of the gantry 44 has been divided into separate slices 107, 108. An irradiation window 109 collimated by the opaque mask 16 (shown in Fig. 1) subtends different portions of the tumor 97 as the tumor 97 is translated through the gantry 44 and the radiation source 12 rotates about the tumor 97. In a helical scan the irradiation window 109 subtends multiple adjacent slices (i.e. it is not slice specific). Referring to Fig. 13(a) at a 0∞ gantry position, the irradiation window 109 may irradiate only slice 107. As the tumor 97 is translated and the radiation source 12 rotated, the irradiation window 109 begins to subtend parts of both slice 107 and adjacent slice 108. Referring to Fig. 13(b), after 90∞ of rotation, the irradiation window 109 may subtend half of slice 107 and half of slice 108. Referring to Fig. 13(c), after 180∞ of rotation, the irradiation window 109 only subtends slice 108. Therefore, helical or non-slice specific data must be developed from the slice sinograms 91 in order to control the compensator leaves 30 during helical irradiation. Although different portions of a tumor may have different cross sections or density distributions, if slice data is generated that corresponds to many thin tumor slices, the changes in the tumor between adjacent slices will be small. Under these circumstances, fluence profiles for adjacent tumor slices directed along identical gantry angles  will be substantially similar. Therefore interpolation between adjacent fluence profiles sharing an identical gantry angles  may be made without appreciably sacrificing irradiation accuracy. To convert the final slice sinograms 91 to a continuous fluence profile "ribbon" 93 for easy use in helical scanning, a helical conversion module 95 (see Fig. 11) may use the following weighted averaging equation to make adequate helical fluence profile approximations: ψ( z 1 , z 2 , j ) = ψ ( z 1 , j ) (1-  j 180 ) + ψ( z 2 , j ) (  j 180 ) where z 1 is the first of two adjacent slices, z 2 is a second adjacent slice,  j is the gantry angle, ψ(z 1 ,z 2 , j ) is the fluence profile of the fan beam 14 directed so as to subtend adjacent portions of tumor slices z 1 and z 2 from gantry angle  j , ψ(z 1 , j ) is the fluence profile from the final slice sinogram 91 corresponding to slice z 1 from gantry angle  j and ψ(Z z , j ) is the fluence profile from the slice sinogram 91 corresponding to slice z 2 from gantry angle  j .  j = 0° when the fan beam subtends only slice z 1 and changes as the gantry rotates until the fan beam subtends only adjacent slice z 2 after 180° of rotation. Therefore, at  j = 0° (see Fig. 13(a)) only ψ(z 1 , j ) will influence ψ(z 1 , z 2 ,  j ) . At  j = 90°, (see Fig. 13(b)) when the translation table has moved the tumor 97 one half of a slice thickness so that one half of the fan beam 14 is directed at slice z 2 , one half of ψ(z 2 , j ) and one half of ψ(z 1 , j ) will influence ψ(z 1 ,z 2 , j ). After the fluence profile ribbon 93 is generated, it is stored in the mass storage system 64 for later use during a therapy session. IV. Operation of The Verification System Prior to a therapy session, the fluence profile ribbon 93 for controlling the compensator 22 is loaded into the compensator control 52. The ribbon 93 consists of a plurality of fluence profiles to be directed at the tumor from a sequence of gantry angles  as the translation table moves through the gantry 44. Each fluence profile consists of desired intensity data for each ray 28 of the fan beam 14. The compensator control 52, directed by the fluence profile ribbon 93 data, drives the leaves 30 into and out of the radiation beam 14 effecting various radiation intensities as described above. Referring to Figs. 3, 4 and 5, as the compensator control 52 drives the leaves 30 between the open and closed states, the trigger member 19 of each leaf 30 is driven in and out of its associated light path 33. When the trigger member 19 blocks the light path 33 (i.e. the leaf 30 is in the closed state), the light detector 15 produces an actual position signal indicating that the leaf 30 is in the closed position. When the trigger member 19 is outside the light path 33 (i.e. the leaf 30 is in the open state), the light detector 15 produces an actual position signal indicating that the leaf 30 is in the open state. An error detector (not shown) realized in software runs on the computer 51 and compares the actual position signals to the desired position signals generated by the compensator control 52 to identify mistakes in leaf 30 movement. If a leaf 30 fails to assume the position indicated by a desired position signal, the compensator control 52 assumes that the actuator 32 failed. Upon failure of an actuator 32, the compensator control 52 directs its signals so as to bypass the primary actuator 32 and begin to direct the backup actuator 35. Thus, a level of redundancy is added to the system wherein failure of a leaf movement means will not require delay in therapy protocol. In the event that both actuators 32, 35 fail, the spring 39 biases the leaf 30 into the closed state, the leaf 30 thus occluding its associated ray 28. In this manner the possibility of a leaf 30 becoming stuck in an open state so that an uncontrolled ray 28 is directed at the patient is reduced. The system can be equipped with an alarm to indicate when one or both of the actuators 32, 35 has failed. However, even upon failure of both actuators 32, 35 associated with the same leaf 30, a therapy session can continue once the leaf 30 is biased into its closed position. A closed leaf 30 poses no danger of radiation overexposure because the associated beam ray 28 is entirely occluded. The deficiency in actual radiation absorbed by the tumor because of leaf 30 failure can be compensated for in later therapy sessions. Referring now to Fig. 14, there is shown a simplified compensator 22 and verification system 71 wherein the compensator 22 has only eight attenuating leaves 30 dividing the fan beam 14 into eight adjacent rays 28. As part of a treatment verification system 59, the first multi-segment ion chamber 47 disposed between the compensator 22 and the patient 17 has eight chamber segments 49, each segment 49 directly within one of the eight rays 28 of the fan beam 14. Each segment 49 produces a measured ray fluence signal 55, or pre-patient fluence signal indicative of the ray fluence encountered thereby. The second multi-segment ion chamber 53 disposed within the fan beam 14 opposing the radiation source 12 on the opposite side of the patient 17 consists of a second group of eight chamber segments 54, each segment 54 intercepting one ray 28c that passes through the patient 17. These chamber segments 54 produce post-patient fluence signals 57 indicative of the fluence of rays 28' that have traversed the thickness of the patient 17. The measured ray fluences 55 are employed by the verification system 71 for two distinct purposes. First, referring to Fig. 15, a comparator module 101 receives from the computer 51 the desired ray fluence to be generated by every leaf 30 of the collimator 22. In addition, the comparator module 101 (see Fig. 15) receives the measured ray fluence 55 generated by the first ion chamber 49. Comparing the desired ray fluence to the measured ray fluence 55, the comparator module 101 generates a difference value 102 for each leaf 30 at each gantry angle . A limit module 104 determines if the difference value 102 is outside of a reasonable limit and dangerous to the patient. If so, the compensator control 52 turns off both the primary 32 and backup 35 actuators for the duration of the therapy session. With the actuators 32, 35 turned off, the biasing spring 39 (see Fig. 4) forces the leaf into the closed position to occlude the ray 28 and eliminate possible radiation danger. If the difference value 102 is minimal or the measured ray fluence 55 is less than the desired ray fluence, the difference can be corrected by adjusting the ray intensity of second ray 28 and a different gantry angle. Referring to Fig. 7, and as described above, the total radiation absorbed by a voxel 74 of the tumor being the summation of radiation directed along many rays 28 toward the voxel 74, discrepancies between desired ray fluence and measured ray fluence 55 can be corrected by adjusting the fluence of the second ray 28 adjacent the first ray. Decreasing fluence of a second ray 28 reduces the quantum of total radiation delivered to a voxel 74. Likewise, increasing fluence of the second ray 28 increases the quantum of total radiation delivered to a voxel 74. In this manner, relatively minor discrepancies between desired and measured ray fluences can be eliminated resulting in more accurate therapy sessions. Second, the measured ray fluence signals 55 are used in conjunction with the post-patient fluence signals 57 to produce tomographic absorption images. By knowing the fluence of each ray 28 entering the patient 17 and the fluence of each ray 28' exiting the patient 17, a simple subtraction calculation generates an absorption value indicating how much radiation is absorbed by the tissue within the patient 17 traversed by the ray 28. By combining all of the rays 28 directed along a gantry angle  while the translation table is in a single position, an absorption profile for that angle  and table orientation can be constructed. Data collected during a helical therapy session in which a fan beam 14 sweeps a helical pattern through a tumor site is not slice specific. Because it is most advantageous to view tomographic images as slices rather than as a helix, the helical data is converted to slice specific data. The computer 51 employs the helical conversion module 95 for a second time to convert the helical data into slice fluence data. Because each tumor slice shares absorption profiles with two adjacent tumor slices (one before and one after), the following weighted averaging equation may be employed: where z 1 is a first tumor slice, z 3 is a second tumor slice, z 2 is a third tumor slice between z 1 and z 3 , ψ(z 2 , j ) is the absorption profile for tumor slice z 2 at gantry angle  j , ψ(z 1 ,z 2 , j ) is the absorption profile detected between adjacent portions of slices z 1 and z 2 at gantry angle  j and ψ(z 2 ,z 3 , j ) is the absorption profile detected between slices z 2 and z 3 at gantry angle  j where  j varies between 0° and 360°. After slice absorption profiles corresponding to the various gantry angles  have been calculated for each tumor slice, tomographic reconstruction techniques may be used to produce a plurality of slice specific tomographic absorption images to be viewed on the display unit 63. Standard isodose curve data used in the art for therapy planning purposes can be used by the computer 51 to establish various levels of radiation absorption along the depth of the tissue traversed by each ray 28. By back-projecting a plurality of absorption profiles (one profile for every gantry angle  at which the beam 14 was directed toward a slice) and simultaneously accounting for the levels of radiation absorption within each ray 28, a tomographic absorption image can be constructed in a manner similar to that used in tomographic x-ray imaging. A radiologist can use the tomographic absorption images to determine radiation dose absorption within slices of the tumor. These images can also be used to develop more accurate diagnostic techniques and to study the specific effects of irradiation on tumor size and longevity. Referring to Fig. 16, a second embodiment of a radiation therapy unit 110 suitable for use with the present invention includes a radiation source 112 producing a generally conical radiation beam 114' emanating from a focal spot 118 and directed towards a patient (not shown). The conical beam 114' is collimated to form a generally planar fan beam 114 centered about a fan beam plane 120. As in the first embodiment, the compensator 122 is centered in the fan beam 114 and about the fan beam plane 120, prior to the.radiation being received by the patient. However, the compensator 122 of the second embodiment includes a rack 124 having an upper level 125 that extends on both sides of the beam 114, well past edges 123 of its thickness 115. Referring to Fig. 17 the upper level 125 of the rack 124 has a set of sleeves 126 that slidably receive a corresponding number of equispaced trapezoidal leaves 128. Each leaf 128 is constructed of a dense, radio-opaque material such as lead, tungsten, cerium, tantalum or a related alloy. The sleeves 126 are constructed of radio-translucent materials and integrally attached to a mounting plate 136 which is fixed relative to the focal spot 118. The mounting plate 136 is constructed of a sturdy radio-opaque material and is positioned outside the fan beam 114 to prevent interference with the fan beam 114. The leaves 128, of the upper level together form an arc of constant radius about the focal spot 118. When the leaves 128 are in the closed position, a leaf gap 130 slightly less than the width of one leaf 128 exists between each two adjacent leaves 128. The leaves 128 and leaf gaps 130 divide the fan beam 114 into a set of adjacent slab like rays 138 at offset angles. The rays 138 that fall on the leaves 128 are attenuated and those directed between the leaves 128 are passed through the upper level 125 unattenuated. Referring to Figs. 16 & 17, the rack 124 also includes a lower level 132 that also extends past the edges of the thickness 115 of the beam 114. Lower level 132 is disposed between the upper level 125 and the patient. Like the upper level 125, the lower level 132 has a second set of sleeves 135 that receive a corresponding second number of equispaced trapezoidal leaves 134. Leaves 134 are constructed of the same material as leaves 128. Each sleeve 135 of the second rack 132 is disposed under a leaf gap 130 in the upper level 125 and is centered within beam 114. Each leaf 134 in the lower level 132 is sized to overlap any flanking leaves 128 above it so that when viewed from the focal spot 118 it eliminates radiation leakage. The overlap, however, of leaves 128 in the upper level 125 and those in the lower level 132 is slight so that very few rays 138 are attenuated by leaves in both racks. The leaves 128, 134 of both levels are substantially the same thickness, or at least thick enough to entirely occlude their associated rays 138 when in the closed position. Referring now to Figs. 16 and 19(a)-(c), the sleeves 126, of both levels 125 are sized to permit each leaf 128 to slide to either side of the fan beam 114 completely out of the path of the fan beam 114 yet to still be guided by the sleeve 126. The sleeve 126, guides the leaf 128, between a "first open state" (Fig. 19(a)), wherein the leaf 128 is positioned to one side of the beam 114 allowing unobstructed passage of one ray 138 of the beam, and a "closed state" (Fig. 19(b)) in which the leaf 128 is positioned in the center of the sleeve 126 occluding one ray 138 of the beam, and a "second open state" (Fig. 19(c)), in which the leaf 128 on the opposite side of the beam 114 allowing unobstructed passage of the ray 138 of the beam. Referring now to Fig. 18, the leaves 128, 134 are supported and guided within the sleeves 126, 135 by guide rails 137 fitted into notches 142 cut along the edges of the leaves 128, 134. The notches 142 allow the guide rails 137 to slidably retain the leaves 128, 134 within the sleeves 126, 135 during motion between the two open and single closed states. Referring again to Fig. 16, each leaf 128 and 134 may be moved rapidly between its open and closed states by means of a corresponding pneumatic cylinder 139 connected to the leaf 128 and 134 by a flexible link 140. The pneumatic cylinders 139 have internal pistons (not shown) that may be moved at high velocity between the ends of the cylinders 139 by means of pressurized air coupled to the cylinders 139 through supply hoses 141. The supply hoses 141 are fed by a compensator control (not shown in Figs. 16 or 17) to be described below. The pneumatic cylinders 139 are capable of applying high forces to the leaves 128, 134 to move them rapidly and independently between the closed and both open states. As best seen in Figs. 16 & 17, two solenoid stops 143 are associated with each leaf 128, 134. One pair of solenoid stops 143 are disposed above each sleeve 126 on the upper level 125, with one stop 143 outside each edge 123 of the beam thickness 115. Similar assemblies 143 are disposed below each sleeve 135 on the lower level 132, one outside each edge 123 of the beam thickness 115. The solenoid stops 143 are used to ensure accurate positioning of the leaves 128, 134 within the fan beam 114 in a manner to be described in more detail below. Referring to Fig. 20, each stop 143 has an electrical solenoid coil 144 which, when energized, attracts a concentrically located armature 145 against the force of a spring 150. A flexible connector 147 connects the armature 145 to a stopping shaft 148 protruding from the stop 143 to be retracted by the armature when the coil is energized. Each solenoid stop 143 is securely fastened to its respective level of the rack 124 by a mounting bracket 146. A flange 149 on the armature 145 holds the spring 150 between itself and the underside of the coil 145 to bias the stopping shaft 148 out from the stop 143 when the coil 144 is de-energized. The flexible connector 147 is longitudinally rigid, but laterally flexible to absorb the torque applied to the stopping shaft 148 by impact with its associated leaf 128, 134 as will be described. The stopping shaft 148 is constructed of a rigid but resilient material that will not bend under the pressure of repeated collisions with a collimator leaf 128, 134. As seen in Fig. 19(b), when the solenoid coil 144 is not energized, the spring 150 biases the armature 145 out from the coil 144 which in turn forces the stopping shaft 148 through a shaft hole 153 in the rack 124, and into the sleeve 126, into the path of leaf movement. When the solenoid coil 144 is energized, the armature 45 retracts into the coil 144 and out of the path of leaf movement. Referring now to Figs. 19(a)-(c), during radiation therapy, the stops 143 remain de-energized most of the time with the stopping shafts 148 protruding into the sleeves and thus locking their associated leaf 128 into one of its three steady state positions, the closed position (19(b)) the first open position (19(a)) and the second open position (19(c)). When a leaf is to move from the first open position (Fig. 19(a)) to the closed position (Fig. 19(b)), the compensator control module 158 first energizes a first solenoid stop 143a which moves a stopping shaft 148a up and out of the sleeve 126 allowing the leaf to move toward the closed position. High pressure air directed to the pneumatic cylinder 139 then pushing the leaf 128 toward the closed position (see Fig. 19(b)). The leaf 128 strikes the second stopping shaft 148b where it is halted in the closed position occluding the entire thickness 115 of the ray 138. Stopping shaft 148(b) ensures the leaf 128 is accurately positioned. The first solenoid stop 143a is then deenergized and the compressively loaded spring 150 expands forcing the stopping shaft 148a down back into the sleeve 126. At this point, the stopping shafts 148a, 148b of both solenoid stops 143a, 143b are within the sleeve 126 and abutting the edges of the leaf 128. With both solenoid stops 143a, 143b de-energized, the leaf 128 is secured in the closed position against perturbing forces. Similarly, to move the leaf 128 into the second open position, the second stop 143b is energized and the stopping shaft 148b is pulled out of the sleeve 126. The leaf 128 is pushed to the second open position (see Fig. 19(c)). The second stop 143b is de-energized and the stopping shaft 148b is forced into the sleeve 126 to lock the leaf 128 out of the beam. As will now be described, by alternating open states from one side of the beam thickness 115 to the other, the characteristic gradient attenuation along the thickness 115 of the ray 138 (i.e. the beam 114) can be eliminated. Referring now to Fig. 21(a), when the leaf 128 is in the closed state, entirely occluding the fan beam thickness 115 the initial cumulative intensity of radiation received by the patient is zero. As the leaf 128 is moving to the first open state (see Fig. 21(b)), the cumulative intensity of the received radiation increases starting with the first areas to be exposed. By the time the leaf 128 reaches the first open state in Fig. 21(c), there is a gradient of received radiation 121 across the beam thickness 115. As the leaf 128 begins to return to the closed state as in Fig. 21(d), the gradient attenuation 121 becomes steeper until it is twice as steep as it was after the first open cycle (see Fig. 21(e)). In the next cycle, the leaf 128 is moved to the second open state as in Fig. 21(f), the exact direction as the previous cycle producing the opposite gradient 121. When the gradients of the two cycles add together, the resulting cumulative intensity 119 of the rays directed across the beam thickness 115 becomes uniform as in Fig. 21(g). Referring again to Figs. 19(a)-(c), it should be noted that alternating movement of a leaf 128, results in a uniform radiation exposure across the thickness of the beam 115 only if the velocity of the leaf 128 is constant as the leaf 128 travels through the beam thickness 115. Without a constant leaf velocity, a parabolic exposure profile across the beam thickness 115 results as the exposure of alternating open cycles is added. While, practically, some acceleration of the leaf 128 as it travels through the beam widths will be acceptable, ideally, most of the acceleration of the leaf 128 will take place while the beam 114 is either fully occluded or entirely unattenuated. Leaf acceleration can be limited to these two conditions by placing the solenoid stop 143 outside the beam thickness 115 and configuring the leaf 128 to be wider than the beam thickness 115 so as to provide for a first 133a and a second 133b acceleration gap on opposite sides of the beam thickness 115. Referring still to Figs. 19(a)-(c), when the leaf 128 moves from the first open state (Fig. 19(a)) to the closed state (Fig. 19(b)), it is accelerated from a stationary position to a constant velocity before the leading edge 129 of the leaf 128 passes the first acceleration gap 133a and enters the beam thickness 115. As the leading edge 129 moves across the beam thickness 115 its velocity is constant until it strikes the second stopping shaft 148b. The leaf 128 is decelerated after the leading edge 129 is within the second acceleration gap 133b and the following edge 131 is within the first acceleration gap 133a. In the closed position (Fig. 19(b)), the leaf 128 occludes the entire beam thickness 115 as well as both acceleration gaps 133a, 133b. Moving from the closed state (Fig. 19(b)), to the second open state (Fig. 19(c)), the leaf 128 is accelerated from its closed stationary position to a constant velocity before the following edge 131 exits the first acceleration gap 133a. Restricting acceleration and deceleration to the period when the leaf edges 129, 131 are within the acceleration gaps is also observed when moving from the second open state to the closed state and from the closed state to the first open state. It has been found that the optimal acceleration gap 133 size is independent of acceleration potential and is one fourth the beam thickness 115 (see Appendix). Therefore, in order for the leaf 128 to occlude both the beam thickness 115 and both acceleration gaps 133 when in the closed state, the leaf 128 should ideally be constructed 1.5 times as long as the beam thickness 115 (see Figs. 19(a)-(c)). Using a leaf 128 that is wider than beam thickness 115. enables the stops 143 to provide a simple leaf deceleration mechanism. When the leaf 128 is moved from its first open state (Fig. 19(a)) to the closed state (Fig. 19(b)), upon impact with the stopping shaft 148b, the leaf 128 may temporarily bounce back off the stopping shaft 148b. Piston pressure is continued to limit "leaf bounce" to positions in which the impacting leaf edge 129 remaining within its associated leaf gap 133b. Thus, leaf deceleration will be limited to acceptable leaf 128 positions. This second compensator embodiment 110 employs the radiation therapy hardware discussed above in connection with Fig. 5, and utilizes the therapy planning software discussed above in connection with Figs. 10 and 11. The above description has been that of a preferred embodiment of the present invention. It will occur to those who practice the art that many modifications may be made without departing from the spirit and scope of the invention. For example, the image reconstructor 60 could produce a tomographic desired fluence image for comparison with the tomographic absorption image to produce a tomographic irradiation error image. The error image could then be used to repeat the helical irradiation process to correct for insufficient irradiation. Clearly the method for planning radiation therapy is not limited to a particular radiation source but may be used with any radiation source which may be decomposed into separately attenuated radiation rays. In addition, a computed tomography system need not be incorporated with the radiation therapy equipment but separate such equipment could be used. The relationship between the terma values and the fluence values may assume a constant density of the patient to eliminate the need for a precise tomographic scan of the irradiated area. It should also be recognized that the pneumatic cylinders 139 in the second embodiment could be constructed to decelerate the leaves 128, 134 and limit the effect of "leaf bounce" and leaf wear upon stopping shaft collisions during leaf movement between states. Referring again to Figs. 19(a)-(c), when leaf 128 is moved from the first open state (Fig. 19(a)) to the closed state (Fig. 19(b)), once the leading edge 129 is through the beam thickness 114, the associated pneumatic cylinder 139 could produce a breaking pulse that slows the leaf 128 before the leading edge 129 impacts the stopping shaft 148b. Similar deceleration pulses could be utilized at the end of each leaf movement to limit leaf and stopping shaft wear. In order to apprise the public of the various embodiments that may fall within the scope of the invention, the following claims are made: Appendix Acceleration and The Optimum Acceleration Gap Treatment planning assumes a number of discrete beams with defined intensities at given gantry angles. In fact, the radiation treatment is done on a continuous basis with constant gantry rotation. For the leaf collimator to produce irradiation approximating a set of discrete beams, the leaf cycle time must be small compared to the time required for the gantry to move between the angles of the theoretical discrete beams. The time for that gantry movement t beam is given by: t beam = T Δ Z LB where T is the total treatment beam-on time, ΔZ is the slice thickness at the isocenter, L is the length of the field (being the sum of many slices) and B is the number of discrete gantry angles for a full rotation (this corresponds to the number of projections angles in CT data acquisition). The minimum beam intensity that can be delivered (other than having the leaf closed for all of time t beam ) depends on the minimum switching time Δt which is the time it takes to open the field and immediately close it. For fastest switching, the closing time should be the same as the opening time. If the velocity of collimator motion is constant then there can be more uniformity in the field as shown in Fig. 21(c). The average beam intensity during switching is ½ of the intensity when the beam is open. A constant velocity means the collimator has to accelerate up to the required velocity before the field is exposed. The minimum switching time is made up of three components: 1. The time it takes to accelerate up to velocity. 2. The time it takes the collimators to completely cross the field. 3. The time to decelerate to zero velocity. With a constant acceleration magnitude, the deceleration time is equal to the acceleration time. The minimum intensity is given by the following equation: Φ min =2 Φ 0 2 t cross + 2 Φ 0 t accel where Φ ˙ o is the fluence rate when the collimator .s open, t cross it the time it takes the collimator to cross the open field and t accel is the time it takes the collimator to accelerate up to a constant velocity (or decelerate to rest). For a constant crossing velocity v and acceleration a, times t cross and t accel are described by simple kinematic ratios. Equation 2 becomes: where Δz is the physical distance of the collimator jaws. This can be expressed in terms of the slice thickness ΔZ projected to the isocenter. Δ z = SCD SAD ΔZ where SCD is the source-to-collimator distance and SAD is the source-to-axis distance. If equation 3 is differentiated with respect to the velocity it is possible to solve for the optimal velocity υ opt : υ opt = Δ za 2 Substituting for the optimal velocity, the minimum intensity Φ min is given by: Φ min = Φ 0 8Δ z a The maximum fluence delivered at a discrete beam position is just the product of the fluence rate Φ ˙ 0 and the time per beam t beam as follows: Φ min = Φ 0 T Δ Z LB The ratio of minimum fluence to maximum fluence is the ratio of Equation 6 to 7: Φ min Φ max = 8Δ z a LB T Δ Z = 8· SCD SAD ·Δ Z · a LB T The optimal switching time t opt is given by: t opt = 2 Δ Z υ opt + 2 υ opt a Substituting in υ opt yields the following: t opt = 3 2Δ Z a Observe that the optimal switching time is the time to completely open the field with a constant velocity, decelerate with a constant acceleration to rest, and Immediately reserve the action to close the field such that there is a minimum amount of fluence emitted during this sequence. The optimal distance through which the constant acceleration takes place before the collimator crosses the beam is given by: z ' opt = υ 2 opt 2 a = Δ z 4 Therefore, referring to Figs. 4(a)-(c), the acceleration gap 133 ( z ' / opt) on either side of the fan beam thickness 115 should be one fourth the thickness of the fan beam 115. Note that the optimal distance z ' / optis independent of acceleration.

Claims: A radiation therapy machine (10,10) having a radiation source (12,112) for producing a radiation beam (14,14) directed toward a patient (17) at a gantry angle (), the beam including a plurality of adjacent rays (28,128), the machine having an attenuating means (22, 122) disposed between the radiation source and the patient, characterised in that the attenuating means is capable of independently controlling the fluence of each ray of the beam, and in that the machine further comprises a verification system comprising: a pre-patient monitor (47)generally disposed between the attenuating means and the patient for determining a pre-patient fluence of each ray of the beam prior to entering the patient; and a post-patient monitor (53)generally disposed opposite the pre-patient monitor with respect to the patient and within the fan beam for determining a post-patient fluence of each ray of the beam exiting the patient. A radiation therapy machine according to claim 1, further comprising an absorption calculator for comparing the pre-patient fluence to the post-patient fluence to produce an absorption value for each ray, the absorption values together providing an absorption profile for the fan beam (14, 114) at a given gantry angle. A radiation therapy machine according to claim 2, further comprising: a reconstruction means (60)receiving absorption profiles from a plurality of gantry angles to create a tomographic absorption image of the patient.

Inventors: HOLMES TIMOTHY, MACKIE THOMAS ROCKWELL, SWERDLOFF STUART, HOLMES, TIMOTHY, MACKIE, THOMAS ROCKWELL, SWERDLOFF, STUART
Applicants: WISCONSIN ALUMNI RES FOUND, WISCONSIN ALUMNI RESEARCH FOUNDATION
Application year: 1994
Main classification: A61N5
IPC codes: G21K5, A61N5, G21K1, C08F234